The present invention relates to a method and an article for improving optical detection and sensitivity. More particularly, the present invention relates to a method and an article for improving optical detection and sensitivity in situations in which emission of fluorescence light is monitored.
Optical detection is used intensively in many fields and for a variety of applications. In many cases, the optical signal emitted by or from a viewed or analyzed object is very low, on the border of detection. Vast efforts are therefore directed at increasing the sensitivity of detection of optical and electro-optical systems, or, in other words, at increasing the ability of optical or electro-optical systems to detect light signals of lesser intensity.
Fluorescence microscopy provides an example. Fluorescence microscopy is one of the most powerful techniques for analyzing tissues and cells [J. S. Ploem (1987) Introduction to Fluorescence Microscopy, Oxford Science Publications, New York]. Unlike bright field microscopy where light is transmitted through an analyzed sample, in fluorescence microscopy, a signal appears only with respect to specific entities that emit light, whereas the background is left dark. This fact makes fluorescence microscopy a very sensitive method for detecting both the existence and distribution of materials in a sample and their quantities. Fluorescence microscopy is therefore one of the most important experimental methods used in light microscopy [Lakowicz (1983) Principles of fluorescence spectroscopy, Plenum Press, New York, London].
Thus, in fluorescence microscopy, an analyzed sample is emitting light, a phenomenon known as fluorescence. The fluorescence light can be native to the analyzed sample, or it can be as a result of an interaction between the analyzed sample and a probe. Some probes are chemicals that fluoresce under certain conditions. For example, probes are known that chemifluoresce differently according to a level of a chemical, e.g., an ion, such as hydrogen or calcium ions, present in the sample or portions thereof. Such probes are therefore useful in determining the concentration and/or distribution of a particular ion in the sample. Other probes include a binding portion and a fluorescent tag. The binding portion can be, for example, a first member of a binding pair, capable of binding a second member of a binding pair present in the sample. The members of a binding pair can be, for example, a ligand that binds a receptor and vice versa, an antibody that binds an antigen and vice versa, a nucleic acid that binds it complement, a substrate, product, inhibitor or analog that binds its enzyme and vice versa, etc. The fluorescent tag is typically a fluorochrome covalently linked to the first member of a binding pair and serves to monitor binding to the second member of the binding pair present in the analyzed sample. Many fluorochromes are presently known each is characterized by a unique absorption spectrum and absorption peak and emission spectrum and peak. Examples of fluorochromes include, fluorescent proteins, such as green, yellow, cian and red fluorescent proteins and smaller chemical compounds such as fluorescein-5-iso-thiocyanate (FITC), rodamine, SpectrumOrange(trademark), SpectrumGreen(trademark), Aqua, Texas-Red, 4xe2x80x2,6-diamidino-2-phenylindole (DAPI), Cy3, Cy5.5. Hundreds of other fluorochromes are known. A partial list of commercially available fluorochromes can be found in the catalog of Molecular Probes. For a detailed review of fluorescent probes see, Mason (editor) (1993) Fluorescent and Luminescent Probes for Biological Activity, Biological Techniques Series, edited by Sattelle, Academic Press Limited, London; Waggoner (1986) Applications of fluorescence in the biomedical sciences, Eds. Taylor et al., New York: Alan R. Liss, Inc. pp. 3-28; and Taylor et al. (1992) The New Vision of Light Microscopy, American Scientist, Vol. 80, pp. 322-335.
In the last few years, further advances have been made in both the detection methods and the fluorochromes. Side by side with the development of fluorochromes that are brighter, more stable and easier to attach to different other compounds, other fluorescent materials have been developed commonly called quantum dots or nanocrystals [see, Bruchez et al (1998) Semiconductor nanocrystals as fluorescent biological labels. Science 281, 2013-2016 and Chan, W. C. et al (1998) Quantum dot bioconjugates for ultrasensitive nonisotopic detection. Science 281, 2016-2018]. These structures are in effect nanosized semiconductors that fluoresce. These structures are far more stable than organic-materials based fluorochromes, and in addition, it is possible to design and manufacture the nanocrystals so that they emit light at a desired spectral range and with a narrower bandwidth.
Improvements were also introduced in the detection of fluorescence. Imaging microscopy employing highly sensitive charge coupled devices (CCD) are used intensively and improve many aspects of detection, including, but not limited to, higher sensitivity, larger number of probes that can be co-detected, accurate quantitative analysis and automation. In addition, confocal microscopy which employs laser scanning mechanisms combined with confocal optics that improves the accuracy in the depth of field is also intensively used [see, Wlison, T. (1990) Confocal Microscopy. Academic Press, London]. These detection methods have broadened the use of fluorescence microscopy.
Fluorescent microscopy was improved to allow detection of different probes simultaneously. A remarked improvement in multicolor fluorochromes is the introduction of combinatorial fluorochromes which are various combinations of few basic fluorochromes [see, Ried et al., (1992) Simultaneous visualization of seven different DNA probes by in situ hybridization using combinatorial fluorescence and digital imaging microscopy. Proc. Natl. Acad. Sci. USA 89, 1388-1392; and, Ried (Jan. 1994) Fluoreszenz in situ Hybridizierung in der genetischen Diagnostik, Faculty of theoretical medicine, Ruprecht-Karls University Heidelberg].
Spectral imaging combined with fluorescence microscopy provides an even better spectral resolution and is presently developed to allow simultaneous detection of several dozens of different combinatorial fluorochromes.
A spectrometer is an apparatus designed to accept light, to separate (disperse) it into its component wavelengths, and measure the lights spectrum, that is the intensity of the light as a function of its wavelength. A spectral imaging device, also referred to herein as xe2x80x9cimaging spectrometerxe2x80x9d is a spectrometer which collects incident light from a scene and measures the spectra of each picture element thereof.
Spectroscopy is a well known analytical tool which has been used for decades in science and industry to characterize materials and processes based on the spectral signatures of chemical constituents therein. The physical basis of spectroscopy is the interaction of light with matter. Traditionally, spectroscopy is the measurement of the light intensity emitted, scattered or reflected from or transmitted through a sample, as a function of wavelength, at high spectral resolution, but without any spatial information.
Spectral imaging, on the other hand, is a combination of high resolution spectroscopy and high resolution imaging (i.e., spatial information). Most of the works so far described in spectral imaging concern either obtaining high spatial resolution information from a biological sample, yet providing only limited spectral information, for example, when high spatial resolution imaging is performed with one or several discrete band-pass filters [See, Andersson-Engels et al. (1990) Proceedings of SPIExe2x80x94Bioimaging and Two-Dimensional Spectroscopy, 1205, pp. 179-189], or alternatively, obtaining high spectral resolution (e.g., a full spectrum), yet limited in spatial resolution to a small number of points of the sample or averaged over the whole sample [See for example, U.S. Pat. No. 4,930,516, to Alfano et al.].
Conceptually, a spectral imaging system comprises (i) a measurement system, and (ii) an analysis software. The measurement system includes all of the optics, electronics and the manner in which the sample is illuminated (e.g., light source selection), the mode of measurement (e.g., fluorescence or transmission), as well as the calibration best suited for extracting the desired results from the measurement. The analysis software includes all of the software and mathematical algorithms necessary to analyze and display important results in a meaningful way.
Spectral imaging has been used for decades in the area of remote sensing to provide important insights in the study of Earth and other planets by identifying characteristic spectral absorption features originating therefrom. However, the high cost, size and configuration of past remote sensing spectral imaging systems (e.g., Landsat, AVIRIS) has limited their use to air and satellite-born applications [See, Maymon and Neeck (1988) Proceedings of SPIExe2x80x94Recent Advances in Sensors, Radiometry and Data Processing for Remote Sensing, 924, pp. 10-22; Dozier (1988) Proceedings of SPIExe2x80x94Recent Advances in Sensors, Radiometry and Data Processing for Remote Sensing, 924, pp. 23-30].
There are three basic types of spectral dispersion methods that might be considered for a spectral imaging system: (i) spectral grating or prism, (ii) spectral filters and (iii) interferometric spectroscopy.
In a grating or prism (i.e., monochromator) based systems, also known as slit-type imaging spectrometers, such as for example the DILOR system: [see, Valisa et al. (Sep. 1995) presentation at the SPIE Conference European Medical Optics Week, BiOS Europe 1995, Barcelona, Spain], only one axis of a CCD (charge coupled device) array detector (the spatial axis) provides real imagery data, while a second (spectral) axis is used for sampling the intensity of the light which is dispersed by the grating or prism as function of wavelength. The system also has a slit in a first focal plane, limiting the field of view at any given time to a line of picture elements. Therefore, a full image can only be obtained after scanning the grating (or prism) or the incoming beam in a direction parallel to the spectral axis of the CCD in a method known in the literature as line scanning. The inability to visualize the two-dimensional image before the whole measurement is completed, makes it impossible to choose, prior to making the measurement, a desired region of interest from within the field of view and/or to optimize the system focus, exposure time, etc. Grating and prism based spectral imaging devices are in use for remote sensing applications, because an airplane (or satellite) flying over the surface of the Earth provides the system with a natural line scanning mechanism.
It should be further noted that slit-type imaging spectrometers have a major disadvantage since most of the picture elements of one frame are not measured at any given time, even though the fore-optics of the instrument actually collects incident light from all of them simultaneously. The result is that either a relatively large measurement time is required to obtain the necessary information with a given signal-to-noise ratio, or the signal-to-noise ratio (sensitivity) is substantially reduced for a given measurement time. Furthermore, slit-type spectral imaging devices require line scanning to collect the necessary information for the whole scene, which may introduce inaccuracies to the results thus obtained.
Filters-based spectral dispersion methods can be further categorized into discrete filters and tunable filters. In these type of imaging spectrometers the spectral image is built by filtering the radiation for all the picture elements of the scene simultaneously at a different wavelength at a time by inserting, in succession, narrow or wider band pass filters in the optical path, or by electronically scanning the bands using acousto-optic tunable filters (AOTF) or liquid-crystal tunable filter (LCTF), see below. Similarly to the slit type imaging spectrometers equipped with a grating or prism as described above, while using conventional narrow-band filters-based spectral dispersion methods, most of the radiation is rejected at any given time. In fact, the measurement of the whole image at a specific wavelength is possible because all the photons outside the instantaneous wavelength being measured are rejected and do not reach the CCD.
Tunable filters, such as AOTFs and LCTFs have no moving parts and can be tuned to any particular wavelength in the spectral range of the device in which they are implemented. One advantage of using tunable filters as a dispersion method for spectral imaging is their random wavelength access; i.e., the ability to measure the intensity of an image at a number of wavelengths, in any desired sequence without the use of filter wheels. However, AOTFs and LCTFs have the disadvantages of (i) limited spectral range (typically, xcex2xcexmin) while all other radiation that falls outside of this spectral range must be blocked, (ii) temperature sensitivity, (iii) poor transmission, (iv) polarization sensitivity, and (v) in the case of AOTFs an effect of shifting the image during wavelength scanning, demanding careful and complicated registration procedures thereafter.
All these types of filter and tunable filters-based systems have not been used successfully and extensively over the years in spectral imaging for any application, because of their limitations in spectral resolution, low sensitivity, and lack of easy-to-use and sophisticated software algorithms for interpretation and display of the data.
A method and apparatus for spectral analysis of images which have advantages in the above respects is disclosed in U.S. Pat. No. 5,539,517, which is incorporated by reference as if fully set forth herein, with the objective to provide a method and apparatus for spectral analysis of images which better utilizes all the information available from the collected incident light of the image to substantially decrease the required frame time and/or to substantially increase the signal-to-noise ratio, as compared to the conventional slit- or filter type imaging spectrometer, and does not involve line scanning. According to this invention, there is provided a method of analyzing an optical image of a scene to determine the spectral intensity of each picture element (i.e., region in the field of view which corresponds to a pixel in an image presenting same) thereof by collecting incident light from the scene; passing the light through an interferometer which outputs modulated light corresponding to a predetermined set of linear combinations of the spectral intensity of the light emitted from each picture element; focusing the light outputted from the interferometer on a detector array, scanning the optical path difference (OPD) generated in the interferometer for all picture elements independently and simultaneously and processing the outputs of the detector array (the interferograms of all picture elements separately) to determine the spectral intensity of each picture element thereof.
This method may be practiced by utilizing various types of interferometers wherein the optical path difference (OPD) is varied to build the interferograms by moving the entire interferometer, an element within the interferometer, or the angle of incidence of the incoming radiation. In all of these cases, when the scanner completes one scan of the interferometer, the interferograms for all picture elements of the scene are completed.
Apparatuses in accordance with the above features differ from the conventional slit- and filter type imaging spectrometers by utilizing an interferometer as described above, therefore not limiting the collected energy with an aperture or slit or limiting the incoming wavelength with narrow band interference or tunable filters, thereby substantially increasing the total throughput of the system. Thus, interferometer-based apparatuses better utilize all the information available from the incident light of the scene to be analyzed, thereby substantially decreasing the measurement time and/or substantially increasing the signal-to-noise ratio (i.e., sensitivity). The sensitivity advantage that interferometric spectroscopy has over the filter and grating or prism methods is known in the art as the multiplex or Fellgett advantage [see, Chamberlain (1979) The principles of interferometric spectroscopy, John Wiley and Sons, pp. 16-18 and p. 263].
In U.S. Pat. No. 5,748,162, which is incorporated by reference as if fully set forth herein, the objective was to provide spectral imaging methods for biological research, medical diagnostics and therapy, which methods can be used to detect spatial organization (i.e., distribution) and to quantify cellular and tissue natural constituents, structures, organelles and administered components such as tagging probes (e.g., fluorescent probes) and drugs using light transmission, reflection, scattering and fluorescence emission strategies, with high spatial and spectral resolutions.
Other uses of the spectral imaging device described in U.S. Pat. No. 5,539,517 are described in the U.S. Patent Nos. 6,088,099 xe2x80x9cMethod for interferometer based spectral imaging of moving objectsxe2x80x9d, 6,075,599 xe2x80x9cOptical device with entrance and exit paths that are stationary under device rotationxe2x80x9d; 6,066,459 xe2x80x9cMethod for simultaneous detection of multiple fluorophores for in situ hybridization and multicolor chromosome painting and bandingxe2x80x9d; 6,055,325 xe2x80x9cColor display of chromosomes or portions of chromosomesxe2x80x9d 5,043,039 xe2x80x9cMethod of and composite for in situ fluorescent hybridizationxe2x80x9d 6,018,587 xe2x80x9cMethod for remote sensing analysis be decorrelation statistical analysis and hardware thereforxe2x80x9d; 6,007,996 xe2x80x9cIn situ method of analyzing cellsxe2x80x9d; 5,995,645 xe2x80x9cMethod of cancer cell detectionxe2x80x9d; 5,991,028 Spectral bio-imaging methods for cell classificationxe2x80x9d; 5,936,731 xe2x80x9cMethod for simultaneous detection of multiple fluorophores for in situ hybridization and chromosome paintingxe2x80x9d; 5,912,165 xe2x80x9cMethod for chromosome classification by decorrelaiton statistical analysis and hardware thereforexe2x80x9d; 5,906,919 xe2x80x9cMethod for chromosomes classificationxe2x80x9d; 5,871,932 xe2x80x9cMethod of and composite for fluorescent in situ hybridizationxe2x80x9d; 5,856,871 xe2x80x9cFilm thickness mapping using interferometric spectral imagingxe2x80x9d; 5,835,214 xe2x80x9cMethod and apparatus for spectral analysis of imagesxe2x80x9d; 5,834,203 xe2x80x9cMethod for classification of pixels into groups according to their spectra using a plurality of wide band filters and hardware thereforexe2x80x9d; 5,817,462 xe2x80x9cMethod for simultaneous detection of multiple fluorophores for in situ hybridization and multicolor chromosome painting and bandingxe2x80x9d; 5,798,262 xe2x80x9cMethod for chromosomes classificationxe2x80x9d; 5,784,162 xe2x80x9cSpectral bio-imaging methods for biological research, medical diagnostics and therapyxe2x80x9d; 5,719,024 xe2x80x9cMethod for chromosome classification by decorrelation statistical analysis and hardware therefore, all of which are incorporated herein by reference.
Spectral imaging systems are potentially useful in all applications in which subtle spectral differences exist between chemical constituents whose spatial distribution and organization within an image are of interest. The measurement can be carried out using virtually any optical system attached to the system described, for example, in U.S. Pat. No. 5,539,517, for example, an upright or inverted microscope, a fluorescence microscope, a macro lens, an endoscope and a fundus camera. Furthermore, any standard experimental method can be used, including light transmission (bright field and dark field), auto-fluorescence and fluorescence of administered probes, etc.
A typical scheme of the optical system of a fluorescent microscope is shown in FIG. 1. The microscope includes an objective lens 10, a light source 12 coupled to a condensor lens 14, a dichroic prism 15 including a exitation filter 16 which serves for illuminating an object 18 with a desired exitation light (solid lines) and an emission filter 17 which serves for transmitting fluorescent light (dashed lines) emitted from object 18 through an ocular lens 20 to an array of detectors 22, such as an eye or a CCD, or in the case of a confocal microscope, to a photomultiplier. In the later case there is also included a scanning mechanism that scans the image picture element by picture element in order to reconstruct an image (not shown).
FIG. 2 shows in more detail the optical path of the rays of light emitted from object 18 which is mounted on a mount 24 (e.g., a microscope slide). As can be seen, only rays that are emitted in the angles defined by a virtual cone 26 (darkened) of angle xcex8 are actually collected by objective lens 10 and are therefore detected by the detector. The angle that spans cone 26 in space is defined by the numerical aperture (NA) of the microscope.
The brightness of an image formed by an objective at a fixed magnification increases with the diameter of the angular aperture (the angle xcex8 of the cone of light collected by the objective). Light rays emanating from the object proceed through air (or an immersion medium such as oil) that lies between the cover glass and the objective front lens. The angular aperture is expressed as the angle between the microscope optical axis and the direction of the most oblique light rays captured by the objective. Mathematically, the numerical aperture is expressed as:                     NA        =                  n          xc3x97          sin          ⁢                      θ            2                                              Eq.  1            
where n is the refractive index of the media positioned between the sample and the objective front lens and xcex8 is the angular aperture, see FIG. 2. The value of n varies between 1.0 for air and about 1.5 for a majority of immersion oils that are used in optical microscopy. The angular aperture is the maximum angle of light rays emanating from the specimen that the objective front lens can capture when the specimen is focused. From the Eq. 1 above, it is obvious that the numerical aperture increases with both angular aperture and the refractive index of the imaging medium.
Theoretically, the highest angular aperture obtainable with a standard microscope objective would be 180 degrees. The sine of 90 degrees is one, which indicates that the numerical aperture is limited not only by the angular aperture, but also by the imaging medium refractive index. A majority of microscope objectives are designed to operate with air (which has a refractive index of 1.0) as the imaging medium between the cover glass and the objective front lens. This yields a theoretical maximum numerical aperture of 1.00, but in actual practice, the highest numerical aperture for a dry objective is about 0.95 (the angular aperture half-angle equals approximately 72 degrees).
Every optical imaging system is diffraction limited by the wave-like nature of light. For an optical system such as a microscope, this limit is well described by the point spread function (PSF). This function describes the distribution of light in the image plane and optical axis as a result of a point light source. This theory is well developed and known, see for example, M. Born and E. Wolf, Principles of Optics 6th edition, Cambridge University Press (Cambridge) and R. H. Webb (1996), Confocal Optical Microscopy, Rep. Prog. Phys. 59 pp. 427-471.
The three-dimensional picture element spread function can be described as follows:
First, lets introduce two generalized variables u and v defined as:                     u        =                                            2              ·              π                        nλ                    ·                      NA            2                    ·          z                                    Eq.  2                                v        =                                            2              ·              π                        λ                    ·          NA          ·          r                                    Eq.  3            
where xcex is the wavelength of the light, z describe the distance from the focal point along the optical axis of the microscope and r is the radial distance from the optical axis in the focal plane. n, as before, is the index of refraction of the material in between the object and the front lens of the objective.
Second, the following Lommel functions is defined:                                           U            n                    ⁡                      (                          u              ,              v                        )                          =                              ∑                          s              =              0                        ∞                    ⁢                                                    (                                  -                  1                                )                            s                        ⁢                                          (                                  u                  v                                )                                            n                +                                  2                  ⁢                  s                                                      ⁢                                          J                                  n                  +                                      2                    ⁢                    s                                                              ⁡                              (                v                )                                                                        Eq.  4                                                      V            n                    ⁡                      (                          u              ,              v                        )                          =                              ∑                          s              =              0                        ∞                    ⁢                                                    (                                  -                  1                                )                            s                        ⁢                                          (                                  v                  u                                )                                            n                +                                  2                  ⁢                  s                                                      ⁢                                          J                                  n                  +                                      2                    ⁢                    s                                                              ⁡                              (                v                )                                                                        Eq.  5            
where the J(x) functions are the Bessel function of different integer order.
Third, the point spread function at (u,v) close to the image point of the microscope is:                               I          ⁡                      (                          u              ,              v                        )                          =                                                                              (                                      2                    u                                    )                                2                            ⁡                              [                                                                            U                      1                      2                                        ⁡                                          (                                              u                        ,                        v                                            )                                                        +                                                            U                      2                      2                                        ⁡                                          (                                              u                        ,                        v                                            )                                                                      ]                                      ·                          I                              0                ⁢                                  xe2x80x83                                                              ⁢                      xe2x80x83                    ⁢          or                                    Eq        .                  xe2x80x83                ⁢        6                                          I          ⁡                      (                          u              ,              v                        )                          =                                            (                              2                u                            )                        2                    ·                      [                          1              +                                                V                  0                  2                                ⁡                                  (                                      u                    ,                    v                                    )                                            +                                                V                  1                  2                                ⁡                                  (                                      u                    ,                    v                                    )                                            -                              2                ⁢                                                      V                    0                                    ⁡                                      (                                          u                      ,                      v                                        )                                                  ⁢                                  cos                  ⁡                                      (                                                                  1                        2                                            ⁢                                              (                                                  u                          +                                                                                    v                              2                                                        u                                                                          )                                                              )                                                              -                              2                ⁢                                                      V                    1                                    ⁡                                      (                                          u                      ,                      v                                        )                                                  ⁢                                  sin                  ⁡                                      (                                                                  1                        2                                            ⁢                                              (                                                  u                          +                                                                                    v                              2                                                        u                                                                          )                                                              )                                                                        ]                    ·                      I            0                                              Eq        .                  xe2x80x83                ⁢        7            
where I0 is the intensity of the point light source in the object plane. Even though the two functions are identical, they are required for performing the numerical calculation for different values of u and v.
This function can be calculated by using numerical evaluations of the Bessel function, see for example: W. H. Press et al (1997) Numerical Recepies in C 2nd edition. Cambridge University Press, New York.
An example of the result is shown in FIG. 3.
Along the optical axis, the PSF becomes rather simple,       I    ⁡          (              u        ,        0            )        =                    (                  sin          ⁢                                    u              4                        /                          u              4                                      )            2        ·                  I        0            .      
A different use of fluorescence detection is in the fields of (i) biological material carrying chips; and (ii) xe2x80x9clab-on-chip devicexe2x80x9d. Biological material carrying chips are small articles including a micro-array of biological materials, such as, but not limited to, cells, phages, bacteria, nucleic acids, proteins, peptides or carbohydrates of a known or unknown nature attached thereto in known positions. A xe2x80x9clab-on-chip devicexe2x80x9d is a more complex chip device which includes an array of micro-vessels and optionally micro-tunnels, micro-valves and micro-pumps and therefore allows for the mixing of different reagents and execution of different reactions in the micro-vessels formed on or in the chip. In both cases, one of the preferred detection methods involves fluorescence light detection.
Other uses of the spectral imaging device described in U.S. Pat. No. 5,539,517 are described in the U.S. Pat. Nos. 6,088,099 xe2x80x9cMethod for interferometer based spectral imaging of moving objectsxe2x80x9d; 6,075,599 xe2x80x9cOptical device with entrance and exit paths that are stationary under device rotationxe2x80x9d; 6,066,459 xe2x80x9cMethod for simultaneous detection of multiple fluorophores for in situ hybridization and multicolor chromosome painting and bandingxe2x80x9d; 6,055,325 xe2x80x9cColor display of chromosomes or portions of chromosomesxe2x80x9d; 5,043,039 xe2x80x9cMethod of and composite for in situ fluorescent hybridizationxe2x80x9d; 6,018,587 xe2x80x9cMethod for remote sensing analysis be decorrelation statistical analysis and hardware thereforxe2x80x9d; 6,007,996 xe2x80x9cIn situ method of analyzing cellsxe2x80x9d; 5,995,645 xe2x80x9cMethod of cancer cell detectionxe2x80x9d; 5,991,028 xe2x80x9cSpectral bio-imaging methods for cell classificationxe2x80x9d; 5,936,731 xe2x80x9cMethod for simultaneous detection of multiple fluorophores for in situ hybridization and chromosome paintingxe2x80x9d; 5,912,165 xe2x80x9cMethod for chromosome classification by decorrelaiton statistical analysis and hardware thereforexe2x80x9d; 5,906,919 xe2x80x9cMethod for chromosomes classificationxe2x80x9d; 5,871,932 xe2x80x9cMethod of and composite for fluorescent in situ hybridizationxe2x80x9d; 5,856,871 xe2x80x9cFilm thickness mapping using interferometric spectral imagingxe2x80x9d; 5,835,214 xe2x80x9cMethod and apparatus for spectral analysis of imagesxe2x80x9d; 5,834,203 xe2x80x9cMethod for classification of pixels into groups according to their spectra using a plurality of wide band filters and hardware thereforexe2x80x9d; 5,817,462 xe2x80x9cMethod for simultaneous detection of multiple fluorophores for in situ hybridization and multicolor chromosome painting and bandingxe2x80x9d; 5,798,262 xe2x80x9cMethod for chromosomes classificationxe2x80x9d; 5,784,162 xe2x80x9cSpectral bio-imaging methods for biological research, medical diagnostics and therapyxe2x80x9d; and 5,719,024 xe2x80x9cMethod for chromosome classification by decorrelation statistical analysis and hardware therefore, all of which are incorporated herein by reference.
Fluorescence detection is acquiring major importance in a variety of technological fields. The desired level of detection, or in other words, the desired level of sensitivity, is increased as samples are becoming smaller and smaller. As an example, in detecting DNA arrays on chip in a process known as hybridization sequencing or in a process of evaluating gene expression, the question that has to be answered is not as simple as a yes/no question. The main issue is the extent to which every sequence is hybridized to as to determine a sequence or an expression level of a gene or genes. The higher the accuracy of measurement is, the more information will result and the more accurate the analysis will be. The detection system is one of the major limiting factors in this sense.
There is thus a great need for, and it would be highly advantageous to have an novel approach for fluorescence detection that will increase the sensitivity by which existing optics can detect fluorescence light.
According to one aspect of the present invention there is provided a light reflecting article comprising a sample carrying article being layered with a light reflecting layer, the light reflecting layer being for allowing an optical collection and detection system to collect both luminescent light emitted from a sample positioned on the light reflecting article in a direction of the optical collection and detection system, as well as luminescent light emitted from the sample in a direction away from the optical collection and detection system and reflected in the direction of the optical collection and detection system via the light reflecting layer, thereby increasing a sensitivity of luminescent light detection.
According to another aspect of the present invention there is provided a method of increasing a sensitivity of an optical collection and detection system in detecting luminescent light emitted from a sample, the method comprising positioning the sample on a light reflecting article, such that the optical collection and detection system collects both luminescent light emitted from a sample positioned on the light reflecting article in a direction of the optical collection and detection system, as well as luminescent light emitted from the sample in a direction away from the optical collection and detection system and reflected in the direction of the optical collection and detection system via the light reflecting article, thereby increasing a sensitivity of luminescent light detection.
According to further features in preferred embodiments of the invention described below, the light reflecting article is a microscopic slide.
According to still further features in the described preferred embodiments the light reflecting article is a chip to which biological material is bound.
According to still further features in the described preferred embodiments the biological material is selected from the group consisting of cells, phages, bacteria, nucleic acids, proteins, peptides and carbohydrates.
According to still further features in the described preferred embodiments the biological material is bound to the chip in array.
According to still further features in the described preferred embodiments the light reflecting article is a lab-on-chip device having a plurality of micro-vessels.
According to still further features in the described preferred embodiments the light reflecting article forms a part of a microscope stage.
According to still further features in the described preferred embodiments the light reflecting layer includes a mirror layer.
According to still further features in the described preferred embodiments the light reflecting layer includes an array of micro-retro-reflectors.
According to still further features in the described preferred embodiments the light reflecting layer includes a stack of thin light reflecting films.
According to still further features in the described preferred embodiments the light reflecting layer provides spectral reflection selected from the group consisting of a single band reflection, a double band reflection, a triple band reflection a multiple band reflection, a longpass reflection and a shortpass reflection.
According to still further features in the described preferred embodiments the optical collection and detection system is selected from the group consisting of a fluorescence microscope and a confocal microscope.
According to still further features in the described preferred embodiments the optical collection and detection system includes a scanable light source.
According to still further features in the described preferred embodiments the scanable light source emits coherent light.
According to still further features in the described preferred embodiments the optical collection and detection system includes an arc lamp.
According to still further features in the described preferred embodiments the optical collection and detection system includes a laser light source.
According to still further features in the described preferred embodiments the luminescent light emitted from the sample originates from aoutofluorescence of the sample.
According to still further features in the described preferred embodiments the luminescent light emitted from the sample originates from a fluorescent probe interacting with the sample.
According to still further features in the described preferred embodiments the luminescent light emitted from the sample originates from a chemifluorescent probe interacting with the sample, the chemifluorescent probe changes an intensity of its emission according to a concentration of a chemical present in the sample.
According to still further features in the described preferred embodiments the fluorescent probe includes a first member of a binding pair conjugated to a fluorescent tag.
According to still further features in the described preferred embodiments the first member of the binding pair is selected from the group consisting of a nucleic acid, a ligand, a receptor, an antigen, an antibody, an enzyme, a substrate, a substrate analog and an inhibitor.
According to still further features in the described preferred embodiments the fluorescent tag is selected from the group consisting of a fluorochrome, a quantum dot, a nanocrystal and a fluorescent protein.
According to still further features in the described preferred embodiments the luminescent light emitted from the sample includes a plurality of emission spectra.
According to still further features in the described preferred embodiments the luminescent light emitted from the sample includes a plurality of emission spectra.
According to still further features in the described preferred embodiments the luminescent light emitted from the sample includes a plurality of emission spectra originating from combinatorial labeling.
The present invention successfully addresses the shortcomings of the presently known configurations by providing an article and method for increasing the sensitivity of fluorescence detection.